Non-invasive analysis, which for purposes of this application includes non-destructive analysis, is a valuable technique for acquiring information about systems or targets without undesirable side effects, such as damaging the system being analyzed. Non-invasive analysis has a broad range of applications including, non-destructive analysis of artifacts for defects, verification of the authenticity of documents, such as, bank notes, bio-metric analysis and bio-medical analysis of living entities. In the case of analyzing living entities, such as human tissue, undesirable side effects of invasive analysis include the risk of infection along with pain and discomfort associated with the invasive process.
In the particular case of measurement of blood glucose levels in diabetic patients, it is highly desirable to measure the blood glucose level frequently and accurately to provide appropriate treatment of the diabetic condition as absence of appropriate treatment can lead to potentially fatal health issues, including kidney failure, heart disease or stroke. A non-invasive method would avoid the pain and risk of infection and provide an opportunity for frequent or continuous measurement.
Non-invasive analysis systems based on several techniques have been proposed. These techniques include: near infrared spectroscopy using both transmission and reflectance; spatially resolved diffuse reflectance; frequency domain reflectance; fluorescence spectroscopy; Polarimetry and Raman spectroscopy.
These techniques are vulnerable to inaccuracies due to issues such as, environmental changes, presence of varying amounts of interfering contamination, skin heterogeneity and variation of location of analysis. These techniques also require considerable processing to de-convolute the required measurement, typically using multi-variety analysis and have typically produced insufficient accuracy and reliability.
A correlation between blood glucose concentration in diabetics and non-invasively measured tissue optical scattering coefficient has been clearly demonstrated, for example in papers published in Optics Letters, Vol. 19, No. 24, Dec. 15, 1994 pages 2062-2064 and OPTICS LETTERS/Vol. 22, No. 3/Feb. 1, 1997 pages 190-192.
This correlation has been proposed as the basis for non-invasive glucose monitoring by involving optical coherence tomography (OCT). This proposed technique depends on a relationship between the slope of the OCT signal and the tissue scattering coefficient or measuring a change in an optical path length with glucose concentration. This approach is described in Proceedings of SPIE, Vol. 4263, pages 83-90 (2001), OPTICS LETTERS/Vol. 26, No. 13/Jul. 1, 2001, Phys. Med. Biol. 48 (2003) 1371-1390 and U.S. Pat. No. 6,725,073 issued Apr. 20, 2004 and U.S. Pat. No. 5,710,630 issued Jan. 20, 1998.
The OCT approach described uses a Super-luminescence diode (SLD) output beam that has a broad bandwidth and short coherence length. OCT analysis involves splitting the SLD output beam into a probe and reference beam (or a first portion of the primary light and a second portion of the primary light). Only a reference beam derived from the same SLD that produced the probe beam can produce interferometrically meaningful signals when combined with the probe related light. The probe beam is applied to the system to be analyzed (the target). Light scattered back from the target is combined with the reference beam to form the measurement signal.
Because of the short coherence length only light that is scattered from a depth within the target such that the total optical path lengths of the probe and reference are equal combine interferometrically. Thus the interferometric signal provides a measurement of the scattering value at a particular depth within the target.
By varying the length of the reference path length, a measurement of the scattering values at various depths can be measured and thus the scattering value as a function of depth can be measured. The profile of the scattering signal as a function of depth is referred to as the OCT signal and the slope or rate of change of the scattering signal as a function of depth is referred to as the slope of the OCT signal.
Accurate correlation between the slope of the OCT signal and the scattering coefficient relies on simplifying assumptions, such as that on average, uniform scattering throughout the tissue. In actual tissue, the scattering elements are cellular membranes and have a highly non-uniform distribution. In a practical device, distortion of the tissue and relative motion between the tissue and the monitoring device also disrupt the correlation between the slope of the OCT signal and the scattering coefficient.
In OCT systems depth scanning is achieved by modifying the relative optical path length of the reference path and the probe path. The relative path length is modified by such techniques as electro-mechanical based technologies, such as galvanometers or moving coils actuators, rapid scanning optical delay lines and rotating polygons. All of these techniques involve moving parts, which have limited scan speeds and present significant alignment and associated signal to noise ratio related problems.
Motion occurring within the duration of a scan can cause significant problems in correct signal detection. If motion occurs within a scan duration, motion related artifacts will be indistinguishable from real signal information in the detected signal, leading to motion related noise or inaccuracies in the measurement of the slope of the OCT signal. Long physical scans, for larger signal differentiation or locating reference areas, increase the severity of motion artifacts and their disruptive effect on correlation between the slope of the OCT signal and glucose concentration.
Non-moving part solutions, include acousto-optic scanning, can be high speed, however such solutions are costly, bulky and have significant thermal control and associated thermal signal to noise ratio related problems. Optical fiber based OCT systems also use piezo electric fiber stretchers. These, however, have polarization rotation related signal to noise ratio problems and also are physically bulky, are expensive and require relatively high voltage control systems.
OCT approaches to measuring glucose concentration typically focus the SLD output beam into the target tissue to perform a depth scan and then repeatedly translate a beam steering mechanism to get multiple depth scans at different locations. The OCT signal is averaged over these multiple scans and the slope of the averaged OCT signal is used to determine the glucose concentration.
The sequential nature of the multiple scan approach further exacerbates the noise related to motion artifacts that is problematic in conventional low speed scanning mechanisms. Using multiple SLDs in parallel to address this problem introduces an un-acceptable cost burden and does not address the fundamental distorting effect of motion on the slope of the OCT signal.
Transient or time varying physical changes in the tissue being analyzed due to, for example, physical compression arising from stress or stress gradients or temperature changes or temperature gradients can cause transient distortions in the depth structure, which can be different for different depth scan locations. Such transient changes can affect the slope of the OCT signal and can do so in a manner that is different for scans at different scan locations.
Furthermore, the non-uniform distribution of scatterers in tissue distorts the slope of the OCT signal within depth scans and between depth scans at different locations (even without motion or transient disruptions). This non-uniform distribution of scatterers in tissue has a disruptive effect on correlation between the slope of the OCT signal and glucose concentration, which is based on only an approximate model. These disruptive interfering influences reduce or randomize the correlation between the slope of the OCT signal and glucose concentration and make using the slope of the OCT signal, or measuring optical path lengths, as the basis for the glucose concentration measurement vulnerable to signal blurring noise.
These aspects cause the conventional approach of using the slope of the OCT signal, or measuring optical path lengths, for monitoring glucose concentration to have significant undesirable signal to noise characteristics and present problems in practical implementations with sufficient accuracy, compactness and robustness for commercially viable and clinically accurate devices. These problematic aspects also present difficulties is using OCT non-invasive analysis for a broad range of applications.
Therefore there is an unmet need for a commercially viable, compact, robust, non-invasive device with sufficient accuracy, precision and repeatability for performing non-invasive analysis such as, measuring analyte concentrations, and in particular measuring glucose concentration in human tissue.